In the present specification, reference is made to the following publications cited for illustrating prior art techniques and conventional implementations of certain procedural measures or partial aspects of excitation and encoding sequences.    [1] P. Mansfield. Multi-planar image formation using NMR spin echoes. J. Phys. C: Solid State Phys. 10: L55-L58 (1977).    [2] H. Fischer, R. Ladebeck. Echo planar imaging image artifacts. In: F. Schmitt, M. K. Stehling, R. Turner, eds. Echo-Planar Imaging: Theory, Technique and Application. Springer, Berlin (1998); pp. 179-200.    [3] M. A. Bernstein, K. F. King, X. J. Zhou. Handbook of MRI Pulse Sequences. Elsevier Academic Press, Burlington, Mass. (2004); pp. 702-739.    [4] M. Doyle, R. Turner, M. Cawley, P. Glover, G. K. Morris, B. Chapman, R. J. Ordidge, R. Coxon, R. E. Coupland, B. S. Worthington, P. Mansfield. Real-time cardiac imaging of adults at video frame rates by magnetic-resonance imaging. Lancet 328(8508): 682 (1986).    [5] B. Chapman, R. Turner, R. J. Ordidge, M. Doyle, M. Cawley, R. Coxon, P. Glover, A. Mansfield. Real-time movie imaging from a single cardiac cycle by NMR. Magn. Reson. Med. 5: 246-254 (1987).    [6] M. S. Cohen, R. M. Weisskoff. Ultra-fast imaging. Magn. Reson. Imaging 9: 1-37 (1991).    [7] J. P. Wansapura, S. K. Holland, R. S. Dunn, W. S. Ball. NMR relaxation times in the human brain at 3.0 Tesla. J. Magn. Reson. Imaging 9: 351-358 (1999).    [8] A. Jesmanowicz, P. A. Bandettini, J. S. Hyde. Single-shot half k-space high-resolution gradient-recalled EPI for fMRI at 3 Tesla. Magn. Reson. Med. 40: 754-762 (1998).    [9] J. S. Hyde, B. B. Biswal, A. Jesmanowicz. High-resolution fMRI using multislice partial k-space GR-EPI with cubic voxels. Magn. Reson. Med. 46: 114-125 (2001).    [10] N. K. Chen, K. Oshio, L. P. Panych. Improved image reconstruction for partial Fourier gradient-echo echo-planar imaging (EPI). Magn. Reson. Med. 59: 916-924 (2008).    [11] P. A. Wielopolski, F. Schmitt, M. K. Stehling. Echo-planar imaging pulse sequences. In: F. Schmitt, M. K. Stehling, R. Turner, eds. Echo-Planar Imaging: Theory, Technique and Application. Springer, Berlin, (1998); pp. 65-139.    [12] R. R. Rzedzian. Method of high speed imaging with improved spatial resolution using partial k-space acquisitions. U.S. Pat. No. 4,767,991 (1988).    [13] M. D. Robson, A. W. Anderson, J. C. Gore. Diffusion-weighted multiple shot echo planar imaging of humans without navigation. Magn. Reson. Med. 38: 82-88 (1997).    [14] S. G. Kim, X. Hu, G. Adriany, K. Ugurbil. Fast interleaved echo-planar imaging with navigator: high resolution anatomic and functional images at 4 Tesla. Magn. Reson. Med. 35: 895-902 (1996).    [15] G. T. Luk Pat, C. H. Meyer, J. M. Pauly, D. G. Nishimura. Reducing flow artifacts in echo-planar imaging. Magn. Reson. Med. 37: 436-447 (1997).    [16] G. T. Luk Pat, C. H. Meyer, J. M. Pauly, D. G. Nishimura. Partial flyback echo-planar imaging. U.S. Pat. No. 5,957,843 (1999).    [17] F. Farzaneh, S. J. Riederer, J. K. Maier, R. Vavrek. View-interleaved EPI on a commercial scanner. Proceedings of the 8th Annual Meeting of the Society of Magnetic Resonance in Medicine, Amsterdam, The Netherlands; p. 832 (1989).    [18] G. C. McKinnon. Ultrafast interleaved gradient-echo-planar imaging on a standard scanner. Magn. Reson. Med. 30: 609-616 (1993).    [19] K. Butts, S. J. Riederer, R. L. Ehman, C. R. Jack. Interleaved echo planar imaging on a standard MRI system. Magn. Reson. Med. 31: 67-72 (1994).    [20] P. Mansfield. Spatial mapping of the chemical shift in NMR. Magn. Reson. Med. 1: 370-386 (1984).    [21] F. Schmitt, P. A. Wielopolski. Echo-planar image reconstruction. In: F. Schmitt, M. K. Stehling, R. Turner, eds. Echo-Planar Imaging Theory, Technique and Application. Springer, Berlin (1998); pp. 141-178.    [22] P. Jezzard. Effects of B0 magnetic field drift on echo planar functional magnetic resonance imaging. In Proceedings of the 4th Annual Meeting of ISMRM, New York, N.Y., USA (1996), p. 1817.    [23] D. C. Noll, C. H. Meyer, J. M. Pauly, D. G. Nishimura, A. Macovski. A homogeneity correction method for magnetic resonance imaging with time-varying gradients. IEEE Trans. Med. Imaging 10: 629-637 (1991).    [24] L. C. Man, J. M. Pauly, A. Macovski. Multifrequency interpolation for fast off-resonance correction. Magn. Reson. Med. 37: 785-792 (1997).    [25] N. K. Chen, A. M. Wyrwicz. Correction for EPI distortions using multi-echo gradient-echo imaging. Magn. Reson. Med. 41: 1206-1213 (1999).    [26] V. J. Schmithorst, B. J. Dardzinski, S. K. Holland. Simultaneous correction of ghost and geometric distortion artifacts in EPI using a multiecho reference scan. IEEE Trans. Med. Imaging. 20: 535-539 (2001).    [27] G. Kruger, A. Kastrup, G. H. Glover. Neuroimaging at 1.5 T and 3.0 T: Comparison of oxygenation-sensitive magnetic resonance imaging. Magn. Reson. Med. 45: 595-604 (2001).    [28] D. N. Guilfoyle, J. Hrabe. Interleaved snapshot echo planar imaging of mouse brain at 7.0 T. NMR Biomed. 19: 108-115 (2006).    [29] G. H. Glover, J. M. Pauly. Projection-reconstruction in MRI. In: D. M. Grant, R. K. Harris (eds.). Encyclopedia of Nuclear Magnetic Resonance. Vol. 8. Wiley, Chichester (1996), p. 3772-3778.    [30] M. K. Stehling, R. J. Ordidge, R. Coxton, P. Mansfield. Inversion-recovery echo-planar imaging (IR-EPI) at 0.5 T. Magn. Reson. Med. 13: 514-517 (1990).    [31] A. Haase. Snapshot FLASH MRI. Applications to T1, T2, and chemical-shift imaging. Magn. Reson. Med. 13: 77-89 (1990).    [32] T. O. Reese, O. Heid, R. M. Weisskoff, V. J. Wedeen. Reduction of eddy-current-induced distortion in diffusion MRI using a twice-refocused spin echo. Magn. Reson. Med. 49: 177-183 (2003).    [33] J. Finsterbusch. Eddy-current compensated diffusion weighting with a single refocusing RF pulse. Magn. Reson. Med. 61: 748-754 (2009).    [34] F. Calamante, D. L. Thomas, G. S. Pell, J. Wiersma, R. Turner. Measuring cerebral blood flow using magnetic resonance imaging techniques. J. Cereb. Blood Flow Metab. 19: 701-735 (1999).    [35] E. L. Barbier, L. Lamalle, M. Décorps. Methodology of brain perfusion imaging. J. Magn. Reson. Imaging 13: 496-520 (2001).
Echo-planar imaging (EPI), first proposed by Mansfield [1] is among the fastest MRI pulse excitation sequences permitting the acquisition of so-called k-space data for a single image in a few tens of milliseconds. The k-space represents the reciprocal lattice space. Due to the high speed, EPI has permitted a number of important imaging techniques, in particular functional brain mapping, perfusion imaging, diffusion imaging, or cardiac imaging. On the other hand, EPI is also prone to a variety of image artifacts (e.g., Nyquist ghosting, chemical shift displacements, distortions and signal voids arising from magnetic susceptibility variations across the object, image blurring due to effective transverse relaxation during the EPI readout), and the resolution is typically limited to more than a millimeter for human applications [2, 3].
Conventional single-shot gradient-echo EPI produces a series of gradient echoes with a bipolar oscillating readout gradient scheme following a radiofrequency (RF) excitation pulse. Each gradient echo is individually phase-encoded by a phase blip [4-6] so that multiple k-space lines are recorded to produce a full (typically two-dimensional, 2D) image in a single-shot fashion. The number of gradient echoes, Netl, is also referred to as the echo-train length, and the interval between two adjacent echoes is referred to as echo spacing, tes [3]. Along the direction of phase-encoding, each echo n is encoded at a different echo time, TE(n) resulting in decaying signal amplitudes according to
                              S          ⁡                      (            n            )                          =                              S            0                    ⁢                      exp            ⁡                          [                              -                                                      TE                    ⁡                                          (                      n                      )                                                                            T                    2                    *                                                              ]                                                          (        1        )            where S0 would be the amplitude of the signal recorded at TE=0, and T2* is the effective transverse relaxation time. Image contrast is predominantly determined by the echo time, TEeff, of the central k-space line. As k-space is sampled from the bottom to the top (alternatively from top to bottom) in conventional blipped EPI methods, this leads to TEeff=TE(Netl/2). Assuming a trapezoidal gradient waveform and data acquisition only during the plateaus of the trapezoids, a gradient system with a maximal amplitude of 45 mT/m and a maximal slew rate of 150 T/m/s, a receiver bandwidth of ±75 kHz (i.e., a dwell time of 6.7 μs), a field of view of 20 cm, and an acquisition matrix of 128×128, the acquisition time for a single echo is Tacq≈0.853 ms, and the minimal echo spacing is tes=1.088 ms. The effective echo time would then be
                              TE          eff                =                                            t              del                        +                                          t                es                            ⁢                                                                    N                    etl                                    -                  1                                2                                              ≈                      70.6            ⁢                                                  ⁢            ms                                              (        2        )            where we have assumed another delay, tdel=1.5 ms, between the RF pulse and the beginning of data acquisition of the first echo. In human brain tissue at 3 T, a typical relaxation time is T2*≈45 ms [7], that is, the image suffers a reduction in the signal-to-noise ratio (SNR) by 79% due to relaxation between excitation of the spin system and detection of the central k-space line carrying the majority of k-space energy. With ramp sampling (i.e., acquiring k-space data during the entire trapezoidal gradient lobe), the minimal echo spacing and effective echo time may be reduced to tes=0.853 ms and TEeff≈55.7 ms, respectively, however, the SNR loss is still 71%. Unless T2* relaxation is the wanted contrast mechanism, as for example in functional brain mapping based on the blood oxygen level dependent (BOLD) effect, this intrinsic signal loss causes problems in the application of EPI for anatomical imaging and seriously limits the implementation of high-resolution EPI techniques.
For achieving short effective echo times in single-shot EPI, partial k-space acquisition is often employed along the phase-encoding direction [6, 8, 9]. It relies on the symmetry properties of k-space, specifically, the fact that the Fourier transform of a real object is Hermitian. In practice, however, additional overscan lines are commonly sampled to correct for unwanted phase shifts and to avoid signal dropout induced by magnetic-susceptibility variations in phase-encoding direction. This loss of signal contributions in k-space cannot be recovered by post-processing strategies [10]. The effective echo time for partial k-space acquisition can be reduced toTEeff=tdel+tes×Nol  (3)where Nol is the number of overscan lines (typically between 8 and 16).
Multi-shot EPI techniques acquire fractions of the k-space data with the echo trains produced by multiple RF excitations [3, 11, 12]. In mosaic EPI, data are acquired as a series of tiles and concatenated prior to image reconstruction [6, 12, 13]. Each tile is acquired within a single echo train, and its position is controlled by adjusting the prephasing gradient area and the polarities of the readout gradient waveform and the phase blips. The major problem is that different phase errors are contained in each tile leading to data inconsistency and image artifacts. The lack of sufficient robustness has so far limited the routine use of mosaic EPI [3, 11]. To correct for such phase errors, overscan lines as in partial k-space acquisition are acquired for the estimation of the phase inconsistencies from the overlapping data. Minimal effective echo times are achieved by reversing the polarity of the phase blips in segmented acquisition schemes to obtain center-out trajectories [14-16]. This also ensures that signal contributions, which are shifted out of the acquisition window for one segment due to variations of the magnetic susceptibility, are recorded during the acquisition of another segment. Signal dropouts due to echo shifting can thus be avoided without overscan lines. Additional navigator lines are measured for correcting intersegment phase and intensity discontinuities [14]. As another consequence of combining segments measured with opposite phase-encoding directions, image distortions due to magnetic field inhomogeneities unfold in opposite directions along the phase-encoding axis. This artifact can be corrected in the Fourier domain if the distribution of the static magnetic field is precisely known, e.g. from a separate field-mapping scan. In previous applications, however, the accuracy of the field-mapping scan was rather low [15, 16] or distortions were even completely ignored [14], which only permits imaging at rather low resolution.
In interleaved multishot EPI [5, 12, 17-19], the phase-encoding amplitudes of the blip areas are increased so that the gap between k-space lines acquired within one echo train is also increased. The k-space lines from the subsequent excitations are placed to fill up the gaps in an interleaved fashion. Due to the shortened echo-train length, this achieves less T2* decay during data acquisition according to Eq. (1) and, hence, increased SNR, reduced image blurring, and an increased effective bandwidth (i.e., reduced chemical-shift displacements and image distortions). Such advantages come at the expense of an increased scan time (linear increase with the number of interleaves) and susceptibility to motion artifacts.